Optical coherence tomography (OCT) is a technology that allows for non-invasive, cross-sectional optical imaging of biological media with hi high sensitivity. OCT is an extension of low-coherence or white-light interferometry, in which a low temporal coherence light source is utilized to obtain precise localization of reflections internal to a probed structure along an optic axis. In OCT, this technique is extended to enable scanning of the probe beam in the direction perpendicular to the optic axis, building up a two-dimensional reflectivity data set, used to create a cross-sectional gray-scale or false-color image of internal tissue backscatter.
OCT has been applied to imaging of biological tissue in vitro and in vivo, although the primary applications of OCT developed to date have been for high resolution imaging of transparent tissues, such as ocular tissues. U.S. patent application Ser. No. 09/040,128, filed Mar. 17, 1998, provides a system and method for substantially increasing the resolution of OCT and also for increasing the information content of OCT images through coherent signal processing of the OCT interferogram data. This system is capable of providing cellular resolution (i.e., in the order of 5 micrometers). Accordingly, OCT can be adapted for high fidelity diagnosis of pre-cancerous lesions, cancer, and other diseases of the skin, bladder, lung, GI tract and reproductive tract using non-invasive medical diagnostics.
In such diagnostic procedures utilizing OCT, it would also be desirable to monitor the flow of blood and/or other fluids, for example, to detect peripheral blood perfusion, to measure patency in small vessels, and to evaluate tissue necrosis. Another significant application would be in retinal perfusion analysis. Monitoring blood-flow in the retina and choroid may have significant applications in diagnosis and monitoring macular degeneration and other retinal diseases. Accordingly, it would be advantageous to combine Doppler flow monitoring with the above micron-scale resolution OCT imaging in tissue.
Wang, et al., xe2x80x9cCharacterization of Fluid Flow Velocity by Optical Doppler Tomography,xe2x80x9d Optics Letters, Vol. 20, No. 11, Jun. 1, 1995, describes an Optical Doppler Tomography system and method which uses optical low coherence reflectrometry in combination with the Doppler effect to measure axial profiles of fluid flow velocity in a sample. A disadvantage of the Wang system is that it does not provide a method to determine direction of flow within the sample and also does not provide a method for generating a two-dimensional color image of the sample indicating the flow velocity and directions within the image.
U.S. Pat. No. 5,549,114 to Petersen, et al. also provides an optical coherence tomography system capable of measuring the Doppler shift of backscattered light from flowing fluid within a sample. However, similar to Wang, et al., Petersen, et al. does not disclose a system or method for displaying the direction of flow within the sample, nor does Petersen, et al. disclose a system or method for providing a two-dimensional color image of the sample which indicates the velocity and directions of flow of fluids within the sample.
The present invention involves an advancement in OCT technology that provides for quantitative measurements of flow in scattering media by implementing systems for increased interferogram acquisition accuracy, coherent signal processing and display. The present invention provides a combination of micron-scale resolution tomographic optical imaging of a tissue sample simultaneously with Doppler flow monitoring of blood and other body fluids within the sample volume.
The present invention utilizes an OCT data acquisition system to obtain precise localization of reflections internal to the sample along the optic axis. Scanning the OCT probe beam in a direction perpendicular to the optic axis builds up a two-dimensional data set comprising a gray-scale cross-sectional image of internal tissue backscatter. The data acquisition system may also utilize a calibration interferometer for providing sub-100 nm accuracy calibration of the reference arm position during interferogram acquisition. In combination with the production of the cross-sectional gray-scale image internal tissue backscatter, the present invention also provides a two-dimensional color data image representing the Doppler flow velocity and directions of moving scatterers within the sample. The method for generating the two-dimensional, color, Doppler image is generally as follows:
For each axial reference arm scan (xe2x80x9cA-scanxe2x80x9d), interferometric data is acquired. This interferometric data corresponds to a cross-correlation measurement of the light returning from the reference arm and the light backscattered from the volume of potentially moving scatterers illuminated by the sample arm probe beam. The acquired interferometric data is first band-pass filtered for noise reduction. Next, the filtered data is coherently demodulated at a frequency corresponding to the Doppler shift induced by the reference mirror of the interferometer to produce an array of in-phase data values vs. time and an array of quadrature data values vs. time. As those of ordinary skill in the art will realize, the time element in these arrays is directly proportional to the reference arm length, and in turn, is directly proportional to the scanning depth into the sample.
Next, a starting time and a time window is selected. The selected starting time corresponds to a starting depth and the selected time window corresponds to a depth-range that is preferably longer than or equal to the coherence length of the light source of the low-coherence interferometer. At the selected starting time, the values of the in-phase array and the quadrature array corresponding to the selected time window are extracted from the in-phase and quadrature arrays and passed into a Fourier transform (FT) circuit or algorithm to obtain a power spectrum (localized Doppler spectrum) for that particular time window (depth range).
From this power spectrum, a central velocity estimate is calculated to obtain an estimate of the mean scatter velocity for that particular time window. In one embodiment, the central velocity estimate is obtained by calculating a centroid for the power spectrum. This velocity estimate will have a sign and a magnitude. Next, one of two colors is assigned for to the velocity estimate, the particular one of two colors assigned being determined according to the sign of the velocity estimate (i.e., blue for positive values and red for negative values). The brightness, intensity or density of the color to assign to the window is determined by the magnitude of the velocity estimate. The resultant color and brightness is applied to an image pixel, or to a plurality of image pixels in a velocity profile (which is an array of image pixels corresponding to the particular A-scan). The particular pixel(s) to which the color and brightness is applied is the pixel(s) corresponding to the particular depth range of the selected window.
Next, at a predetermined point past the starting time, another time window of in-phase and quadrature data is extracted from the in-phase an time window, a velocity estimate is calculated as described above, a color is assigned for this velocity estimate, and the color is applied to the velocity profile corresponding to the particular depth range of this next time window. New windows will thereafter be repeatedly extracted and processed to generate a complete velocity profile for the particular A-scan. In one embodiment of the invention, this complete velocity profile is a one-dimensional array of colorized image pixels corresponding to the lateral position of the sample probe.
Preferably, multiple A-scans are performed for a sample, each A-scan corresponding to a particular lateral position of the sample arm with respect to the sample. Alternatively, for angular scanning implementations of OCT, subsequent A-scans are arranged into a radial array. It is to be understood that while the present invention describes multiple A-scans as being displaced laterally, it is within the scope of the invention to utilize the same signal processing algorithms described herein with respect to these acquisition and display methods. The above process of generating a velocity profile is performed for each A-scan taken. The processing can be performed either in real-time or after all of the A-scans have been taken and stored in memory. Thereafter, the velocity profiles are aligned side-by-side to create a two-dimensional colorized Doppler image indicating the direction and velocity of moving scatterers within the sample. This two-dimensional colorized Doppler image may then be merged with, or superimposed onto the gray-scale tissue backscatter image to create a complete image of the sample showing internal tissue scatterers as well as the velocity and direction of moving scatterers within the sample.
The time windows extracted from the real array and imaginary arrays, described above, can be either non-overlapping (adjacent) or overlapping time windows. If non-overlapping windows are used, the color determined by the above process would be applied to the entire time window (i.e., if the time windows correspond to a sample depth that will be approximately 10 pixels deep in the final image, the color determined by the above process would be applied to all 10 pixels in the colorized Doppler image). But if overlapping time windows are used, the velocity value determined by the above process can be applied to a time (depth) range equal to the amount that the time windows do not overlap. For example, if the windows overlap entirely except for a sample depth corresponding to a depth of approximately one pixel in the final image, the color determined by the above process would be applied to only a single pixel in the colorized Doppler image. Thus, while more processing would be required to process the entire A-scan data, the overlapping time windows would provide a much more precise Doppler image than the non-overlapped time windows.
An alternate scheme for color coding the velocity information assigns a specific color or shade to each velocity and encodes into brightness, the magnitude of the signal or equivalently, the reflectivity of the tissue at the site. This alternate color coding scheme may offer some advantages in image rendering speed and ease of interpretation. In this scheme, a blood vessel would appear as a series of concentric rings, whose radius is smallest at the vessel""s center.
Prior to merging the two-dimensional colorized Doppler image with the gray-scale image the user may be prompted to select both positive and negative, minimum threshold and maximum saturation velocities. These threshold and saturation velocities are then applied to the velocity array or to the velocity profile to provide an upper and lower bound to the velocity display color scale which allows all desired velocities to be displayed. Velocities which are larger than the maximum positive and negative saturation velocities, respectively, may be rendered the same color (brightness, shade, etc.) as the positive or negative maximum saturation velocities themselves, or else may be rendered in some other distinctive color (brightness, shade, ect.) in order to indicate that the velocity color scale was saturated at that location. Velocities which are smaller than the positive and negative threshold velocities are rendered as transparent to indicate that flow was not detected in that image region, and to allow for visualization of the underlying gray-scale magnitude reflectivity image.
It has been shown, mathematically, that a better velocity estimate can be obtained for a particular position within the sample (i.e., within a particular time window and lateral position) by taking multiple A-scans for that lateral position and then averaging the velocity estimates together. Accordingly, in one embodiment of the invention, multiple A-scans are taken for each lateral position of the sample arm, and the velocity estimates calculated for each A-scan taken are averaged together to produce a more accurate velocity estimate for that particular lateral position of the sample arm. Alternatively, better velocity estimates may also be obtained by averaging estimated windowed spectra obtained from multiple A-scans for each time window and lateral position, followed by calculating the centroids of the averaged spectra.
It is also known that the spatial resolution of the velocity estimate is limited by the window size, the larger the windowxe2x80x94the lower the spatial resolution. But velocity resolution is inversely related to the window size, the larger the windowxe2x80x94the better the velocity resolution. Two alternate embodiments of the invention, however, overcome this trade-off and provide for simultaneous high spatial resolution and high velocity resolution. This is done, in a general sense, by slowing down the time for taking the A-scan.
In a first alternate embodiment, for each position of the reference arm, instead of one set of in-phase and quadrature data taken, multiple sets of in-phase and quadrature data are taken. Thus, instead of calculating a velocity estimate for a window pertaining to multiple positions of the reference arm, the velocity estimate is calculated for a block of data pertaining substantially to a single position (depth) of the reference arm. With this alternate embodiment, it is not necessary to take the multiple sets of data at a single position of the reference arm, so long as all the readings are taken within a range of reference arm movement less than the coherence length of the interferometer""s light source.
In a second alternate embodiment, instead of one A-scan being taken for each lateral position of the sample arm, multiple A-scans are taken at each lateral position of the sample arm. The in-phase and quadrature data are then arranged into a two-dimensional matrix by aligning the one-dimensional (vertical) matrices for each A-scan taken side-by-side. Thus, instead of calculating a velocity estimate for a window pertaining to multiple positions of the reference arm (a vertical xe2x80x9ccolumnxe2x80x9d window), the velocity estimate is calculated for a single position (a horizontal xe2x80x9crowxe2x80x9d window) of the reference arm. The step of taking multiple A-scans for each lateral position of the sample arm may be accomplished by: (a) incrementally advancing the sample arm laterally and taking the multiple A-scans for each incremental advance of the sample arm; (b) laterally sweeping the sample arm over the sample a multitude of times, taking one A-scan for each lateral position of the sample arm; or (c) some combination of (a) and (b).
The above steps for converting the interferogram data into power spectrum data may be performed by a bank of narrow-band band-pass filters (NBPF), where each NBPF passes a particular frequency along the power spectrum frequency scale. The outputs of each NBPF may be input directly into the centroid calculation. This method eliminates the need for the short-time Fourier transform circuit/algorithm and, may also eliminate the need for the coherent demodulation circuit/algorithm. Thus, this method provides faster and cheaper signal processing.
The above steps for converting the interferogram data into power spectrum data may also be performed by a bank of coherent demodulators and a corresponding bank of low-pass filters, where each demodulator demodulates the data at a particular frequency along the power spectrum frequency scale. This approach offers several advantages for the acquisition of local spectral information. First, the output of each of the detection channels is nearly instantaneous, allowing for substantially real-time implementation of CDOCT. Further, the electronic components are readily available.
This latter approach may also be accomplished using a synthesized multi-frequency demodulating waveform, a mixer, and a low-pass filter. In waveform is synthesized as a sum of a desired number of sinusoids at a range of frequencies. These frequencies are mixed in a single mixer with the interferogram data, and the resulting mixed signal is then low-pass filtered. While this alternate embodiment represents a substantial savings in the number of components required, it does not provide estimation of the complete frequency spectrum within the analysis region.
Finally, the present invention provides two implementations of scatterer velocity estimation and one implementation of turbulence estimation using an auto-correlation method, which is based on a mathematical derivation that estimates the central velocity and turbulence of the fluid flow in the sample as being related to the first and second moments of the Doppler-shifted frequency power spectrum. The first utilizes a complex demodulated detector output to obtain both the centroid velocity and an estimate of the power spectrum variance. The second is a much simpler version and provides an estimate of the central velocity only.
Accordingly, it is an object of the present invention to provide a method for generating a two- or three-dimensional color image of a sample indicating the flow velocity and directions within the image using an OCT acquisition system.
It is a further object of the present invention to provide a method for generating a velocity-indicating, tomographic image of a sample in an optical coherence tomography system including the steps of (a) acquiring cross-correlation data from the interferometer; (b) generating a grayscale image from the cross-correlation data indicative of a depth-dependent positions of scatterers in the sample; (c) processing the cross-correlation data to produce a velocity value and location of a moving scatterer in the sample; (d) assigning a color to the velocity value; and (f) merging the color into the grayscale image, at a point in the grayscale image indicative of the moving scatterer""s location, to produce a velocity-indicating, tomographic image.
These and other objects and advantages of the present invention will be apparent from the following description, the appended claims and the attached drawings.